Method and apparatus for monitoring glucose levels in a biological tissue

ABSTRACT

In accordance with the invention, a low coherence interferometer is used to non-invasively monitor the concentration of glucose in blood by shining a light over a surface area of human or animal tissue, continuously scanning the light over a two dimensional area of the surface, collecting the reflected light from within the tissue and constructively interfering this reflected light with light reflected along a reference path to scan the tissue in depth. Since the reflection spectrum is sensitive to glucose concentration at particular wavelengths, measurement and analysis of the reflected light provides a measure of the level of glucose in the blood. The measurement of glucose is taken from multiple depths within blood-profused tissue, and sensitivity is preferably enhanced by the use of multiple wavelengths. Noise or speckle associated with this technique is minimized by continuously scanning the illuminated tissue in area and depth.

FIELD OF THE INVENTION

This invention pertains to methods and apparatus for monitoring theblood glucose level in biological tissues. It is especially useful fornon-invasive monitoring of blood glucose in human or animal diabetics.Specifically, it relates to methods and apparatus for non-invasivemonitoring of blood glucose using optical coherence interferometry withcontinuous area scanning and simultaneous depth scanning to reduce theeffect of speckle (noise).

BACKGROUND OF THE INVENTION

Monitoring of blood glucose (blood sugar) concentration levels has longbeen critical to the treatment of diabetes in humans. Current bloodglucose monitors involve a chemical reaction between blood serum and atest strip, requiring an invasive extraction of blood via a lancet orpinprick. Small handheld monitors have been developed to enable apatient to perform this procedure anywhere, at any time. But theinconvenience of this procedure—specifically the blood extraction andthe use and disposition of test strips—has led to a low level ofcompliance. Such low compliance can lead to serious medicalcomplications. Thus, a non-invasive method for monitoring blood glucoseis needed.

Studies have shown that optical methods can detect small changes inbiological tissue scattering related to changes in levels of bloodsugar. Although highly complex, a first order approximation ofmonochromatic light scattered by biological tissue can be described bythe following simplified equation:I _(R) =I _(O)exp[−(μ_(a)+μ_(S))L]where I_(R) is the intensity of light reflected from the skin, I_(O) isthe intensity of the light illuminating the skin, μ_(a) is theabsorption coefficient of the skin at the specific wavelength of light,μ_(S) is the scatter coefficient of the skin at the specific wavelengthof light, and L is the total path traversed by the light. From thisrelationship it can be seen that the intensity of the light decaysexponentially as either the absorption or the scattering of the tissueincreases.

It is well established that there is a difference in the index ofrefraction between blood serum/interstitial fluid (blood/IF) andmembranes of cells such as blood cells and skin cells. (See, R. C.Weast, ed., CRC Handbook of Chemistry and Physics, 70th ed., (CRCCleveland, OH 1989)). This difference can produce characteristicscattering of transmitted light. Glucose, in its varying forms, is amajor constituent of blood/IF. The variation of glucose levels inblood/IF changes its refractive index and thus, the characteristicscattering from blood-profused tissue. In the near infrared wavelengthrange (NIR), blood glucose changes the scattering coefficient more thanit changes the absorption coefficient. Thus, the optical scattering ofthe blood/IF and cell mixture varies as the blood glucose level changes.Accordingly, an optical method has potential for non-invasivemeasurement of blood glucose concentration.

Non-invasive optical techniques being explored for blood glucoseapplication include polarimetry, Raman spectroscopy, near-infraredabsorption, scattering spectroscopy, photoacoustics and optoacoustics.Despite significant efforts, these techniques have shortcomings such aslow sensitivity, low accuracy (less than current invasive home monitors)and insufficient specificity of glucose concentration measurement withinthe relevant physiological range (4-30 mM or 72-540 mg/dL). Accordingly,there is a need for an improved method to non-invasively monitorglucose.

Optical coherence tomography, or OCT, is an optical imaging techniqueusing light waves that produces high resolution imagery of biologicaltissue. OCT creates its images by interferometrically scanning in deptha linear succession of spots, and measuring absorption and/or scatteringat different depths in each successive spot. The data is then processedto present an image of the linear cross section. It has been proposedthat OCT might be useful in measuring blood glucose.

There are, however, major drawbacks to the use of OCT for glucosemonitoring. First, the OCT process requires lengthy scanning to reduceoptical noise (“speckle”). Speckle arises from wavefront distortion,when coherent light scatters from tissue. OCT seeks to minimize speckleby averaging it out over many measurements. However this approach in OCTrequires a time period impractically long for a home monitor, and eventhen speckle in OCT remains problematic for achieving a sufficientlyaccurate measurement of glucose level.

A second drawback of OCT is that it requires complex processing to forman image and even further processing to analyze the image data todetermine glucose levels.

A third drawback is that OCT requires expensive, bulky, precisionequipment neither suitable for transport or for use outside thelaboratory. Accordingly, there is a need for an improved methods andapparatus for non-invasive blood glucose monitoring.

SUMMARY OF THE INVENTION

In accordance with the invention, the blood glucose concentration withina biological tissue is monitored by providing light having scatteringproperties sensitive to the glucose concentration within the tissue andcontinuously scanning the light over a two dimensional area of thetissue while, at the same time, interferometrically scanning the tissuein depth. The light reflected from the scanned tissue is collected andanalyzed to determine the concentration of glucose in the tissue. In anadvantageous embodiment light from one or more sources is split into asample beam and a reference beam. The sample beam is continuouslyscanned over the surface and the phase of the reference beam is variedand interfered with the reflected light to effect interferometric depthscanning. In a preferred embodiment, the light provided is composed ofat least two different wavelengths having measurably differentscattering properties for glucose-containing tissue or biologicalindicators of such tissue.

An apparatus for measuring the blood glucose level in biological tissuecomprises one or more light sources to provide light. Optical fiber orlens-directed paths direct the light onto the tissue, an area scannercontinuously scans the light over a two-dimensional area of the tissueand an interferometer effectively scans the tissue in depth. Theinterferometer also collects, analyzes and measures the light reflectedwithin the tissue. A processor responsive to the light measurements thencalculates the glucose level of blood-profused regions of the tissue.Advantageously, the apparatus uses low coherence light sources (lightemitting diodes (LEDs) or super luminescent diodes (SLEDs), a lowcoherence interferometer (LCI), and beam focusing optics. The continuousscanning over a two-dimensional area and the depth scanning reduce noiseor speckle and optimize the amount of blood-profused tissue scanned.

BRIEF DESCRIPTION OF THE FIGURES

The advantages, nature and various additional features of the inventionwill appear more fully upon consideration of the illustrativeembodiments now to be described in detail in connection with theaccompanying drawings. In the drawings:

FIGS. 1 a and 1 b illustrate the anatomy of the skin and the resultingOCT signal;

FIGS. 2 a and 2 b graphically represent the data collected from an OCTsignal and blood measurements and the correlation between the two;

FIG. 3 is a schematic diagram illustrating a method of blood glucosemonitoring in accordance with the invention;

FIG. 4 schematically illustrates an apparatus useful for practicing themethod of FIG. 3;

It is to be understood that the drawings are for the purpose ofillustrating the concepts of the invention, and except for the graphs,are not to scale.

DETAILED DESCRIPTION OF THE INVENTION

FIG. 1 a is a schematic cross section of the skin showing the skinsurface 1, the epidermis layer 2 and the dermis layer 3. Variousadditional structures are labeled. Some structures, such as hair and itsassociated follicles, scatter light in a manner unrelated to glucoseconcentration. Other regions, such as regions in the dermis near thecapillaries are profused with blood and IF (blood profused tissues) andlight scattered from these regions is highly correlated with glucoseconcentration.

FIG. 1 b graphically illustrates light scattered within the skin as afunction of depth. Region 1 of the curve corresponds to light reflectedfrom the skin surface. Region 2 shows light scattered within theepidermis, and Region 3 is light scattered from the dermis. As we willshow, slope of the curve at depths where the tissue is profused withblood (e.g. near the capillaries) is highly correlated with bloodglucose levels.

FIG. 2 a graphically plots the slope of the optical signal reflected byblood-profused tissue (the continuous line) and the measured bloodglucose concentration (rectangular dots) as a function of time after ahuman subject ingests a glucose drink. As can be seen, the slope of thereflected signal closely follows the glucose concentration.

FIG. 2 b illustrates the best-fit linear relation between the slope ofthe reflected optical signal in FIG. 2 a with the measured blood glucoseconcentration. As can be seen, the slope of the reflected light inregions of blood-profused tissue is highly correlated with glucoseconcentration. These data show a 0.95 correlation.

Applicant has determined that the information contained in scatteredoptical signals can be used to measure glucose concentration withoutcreating and analyzing images. At a wavelength of around 1.3 μm, thescattering coefficient, μ_(S), is several times larger than theabsorption coefficient, μ_(a), and a simple linear fit to thelogarithmic data will yield the scattering coefficients within thedermis layer. For blood-profused regions, this scattering coefficientcan be related to glucose concentration changes within the blood.

A non-invasive glucose monitor ideally overcomes three problems. First,the speckle (noise) should be overcome within a reasonable test time.Second, the effect of tissues other than blood-profused tissues shouldbe minimized, and third, as much blood-profused tissue as possibleshould be analyzed. Applicant addresses these problems by continuouslyscanning a two-dimensional area of the tissue to be monitored(preferably the skin) while at the same time interferometricallyscanning to different depths. The area scanning optionally combined withthe multiple level depth scanning permits measurement of many differentlocalized areas of blood-profused tissue, minimizing the effect ofspeckle. Preferably, the light comprises at least two differentwavelengths of light having measurably different absorption andscattering properties for glucose, blood, or other biological indicatorsof blood-profused tissues. The two different wavelengths permitidentification of which measurements are of blood-profused tissues.

In some respects, the invention can be thought of as a modification ofOCT from an imaging technique to a non-imaging technique. Since no imageis formed, the collection optics and data processing are greatlysimplified. In addition, the surface scanning is changed from stepwiselinear scanning to continuous or near continuous two dimensional areascanning to more rapidly and effectively reduce speckle.

Adapting an OCT system to blood glucose sensing, instead of imaging,provides non-invasive glucose concentration monitoring of a human oranimal subject, because the scattered optical signal data can becollected and processed more easily and the resulting data can becorrelated linearly to the level of glucose. Additionally, continuousscanning of the incident light beam over a two-dimensional area on theskin surface, instead of the conventional one-dimensional straight line,markedly improves. signal stability while dramatically reducing thenoise associated with tissue inhomogeneity. Advantageously an area inthe general shape of a circle is scanned.

More specifically, the invention uses interferometry, preferably lowcoherence interferometry (LCI), to measure the glucose concentration ofhuman or animal blood-profused tissue. LCI can be effected withconvenient low coherence light sources and provides advantageously smallvolumes of constructive interference (“regions of optical interaction”)that can be used to localize readings to blood profused tissue. Beamfocusing can be used to further localize the regions of opticalinteraction. LCI uses a standard interferometer illuminated by lowcoherence light sources. The interferometer can be any one of thestandard forms, e.g. Michaelson, Mach-Zehnder, or the like. The lightsources can be LEDs or SLEDs.

The invention is advantageously directed to measure the blood glucoseconcentration from tissue located in the dermis layer of the skin (3 ofFIGS. 1 a and 1 b). In this layer, the rate at which the light intensitydecays is a function of the scattering and absorption coefficients,μ_(S) and μ_(a), respectively. The location of the blood profused tissuein the skin varies from subject to subject. It is often found just belowthe dermis/epidermis junction (100-350 microns depth) and in a deepregion below the skin surface (greater then 600 microns).

With low coherence sources, the interferogram is generated over a smallvolume whose position in the depth of the object can be determined viathe phase of a reference beam. This, in turn, can be controlled by aphase shifter, such as a movable reference reflector, preferably amirror. Thus, a high degree of localization of the measured scatteringphenomena can be achieved. For example, for a typical light emittingdiode (LED) operating at a 1.3 μm wavelength, a depth resolution of 10μm is easily achieved in biological tissues. Typically in biologicaltissues, the scatter of the light occurs at the interface between thecell membrane and the fluid that surrounds the cell (i.e. blood orinterstitial fluid). Measuring this scatter permits the glucose level tobe determined, as by the linear fit shown in FIG. 2 b.

FIG. 3 is a schematic block diagram of the method of measuring the bloodglucose concentration on a human or animal subject. The first step,shown in Block A, is to provide light having scattering absorption orproperties sensitive to glucose concentration within the tissue.Preferably the light provided comprises at least two differentwavelengths. By different wavelengths is meant that the wavelengthsshould be sufficiently different that they have measurably differentabsorption and scatter properties for different levels of glucose and/orindicator components such as blood. Typically, the light is providedfrom multiple single wavelength sources, such as low coherencesuperluminescent diodes (SLEDs) at wavelengths in the red/near infraredrange (RNIR). Alternatively, the light can be provided from a singlebroadband source appropriately notch filtered. Both wavelengths of lightare advantageously directed in a single beam.

The next step, shown in Block B, is to split the single beam of lightinto a reference beam and a sample beam. The reference beam travels inan adjustable phase path denoted as the reference beam path (referencearm), and the sample beam travels in a sample beam path (sample arm)where it is directed onto the tissue to be monitored, e.g. the skin of ahuman diabetic. The light in the reference beam is directed over anadjustable phase path and will subsequently be interfered with samplelight reflected from within the tissue.

In the third step, Block C, the sample beam is continuously or nearcontinuously scanned over a two-dimensional area of the tissue while, atthe same time, being interferometrically scanned in depth. Block D showsvarying the phase (path length) of the reference beam so that light fromthe reference beam constructively interferes with reflected sample lightfrom successively different depths of tissue. Block E shows thereflected light collected and interfered with the reference beam. As theinterferometer sweeps in depth, the surface scan is also sweepingcontinuously. This “smears” out the scan and reduces the effect ofspeckle.

The next steps, Blocks F, G, and H are to process the resulting data tocalculate glucose concentration. In essence, this is achieved bycomputing the scattering coefficient of glucose-containing tissue. BlockF indicates the scanning data is input into a digital processor. BlockG, which is optional, but advantageous, is to identify those scatteringmeasurements that are from blood-profused tissue (in or near bloodvessels). Such identification can be accomplished, for example, byproviding light of two different wavelengths, at least one of whichscatters from blood profused tissues in a characteristic manner.Finally, in Block H, the scattering coefficient of the glucosecontaining tissue is calculated, and the correlated glucose level inblood is determined.

FIG. 4 schematically shows advantageous apparatus 400 for practicingmethod of FIG. 3. The apparatus 400 comprises a fiber optics based lowcoherence interferometer (LCI). A 2×2 fiber optic. splitter 401 formsthe basic interferometer. An optical input from light sources 406 issplit between a sample beam 402 and a reference beam 404. Sample lightin beam 402 is continuously scanned across a sample surface by scanner408. Preferably, the end of the sample beam 402 can contain imagingoptics 403 to tailor the spot size according to the area tissue beingmeasured. Reference beam 404 is varied or adjusted in phase as by amoveable mirror 405 which can be vibrated or oscillated to scan depth.Reflected signals from beams 402 and 404 interfere and are presented tophotodetector 407 for measurement. Advantageously, imaging optics 403can provide high coupling efficiency between the optical system and thetissue.

The tissue volume with which the light interacts (referred to as theinteraction volume) is determined by the spot size of the imaging optics(surface area) and the coherence length of the light (depth) Thereference beam 404 has a scanning reflector 405 (such as a mirror). Thereflector 405 of the interferometer determines the phase shift appliedto the reference beam 404 and thus, which reflected light from thereference beam 404 will constructively interfere with the reflectedsample beam 403. The differences in phase of the beams determines thedepth from which scattered light will be measured. This can permit afixed depth, adjustable depth, or a scan of multiple depths within thetissue. LCI is thus sensitive to the intensity of the reflected lightlocalized in a small volume of tissue. Determination of the depth andinteraction volume permits a more accurate selection of regions ofblood-profused tissue beneath the skin.

A photodetector 407 (such as a photodiode) can be used to measure theinterference of the light from both the sample beam 402 and thereference beam 404. One or more photodetectors 407 may be used alongwith optical filters (not shown) designed for each of the differentwavelength light sources 406 used in the measurement.

Preferably, the imaging optics 403 are beam focusing optics to reducethe beam cross section so as to minimize the region of opticalinteraction with the tissue. The use of these optics will enhance theselectivity of the signal while also reducing the effect of speckle.

Light passing through turbid biological tissue is subject to wavefrontdistortion that produces coherent noise or “speckle”. The effect ofspeckle can be reduced by taking multiple scans from different locationson the tissue and then averaging these scans. This solution isimpractical for the typical OCT imaging system, because the vast numberof scans needed to reduce speckle would take too long and would producea severe loss in the resolution of the image. However, for the presentinvention, the collection optics can be simpler. The present non-imagingsystem presents a practical solution to reducing coherent noise. Notonly does the speckle effect significantly decrease, but the non-imagingsystem can continuously scan a two-dimensional area of tissue instead ofbeing limited to a single scanning line. Area scans reduce speckle dueto the diversity of tissue regions encompassed in the scan. They alsomaximize the coverage of blood-profused tissue. Thus, coherent noise isalso further reduced.

An alternative solution is to use parallel optical processing wheremultiple spots on the subject tissue are measured together to create“boiling” speckle. Boiling speckle occurs where the sub-spot speckle ischanging so quickly that the observed speckle is averaged out by thehuman eye, or the integration time of the optical receiver. Thisinventive system may be modified to create boiling speckle by replacingthe scanner 408 with either a lenslet array or a diffractive opticalelement (DOE). If the lenslet or DOE is rapidly translated or rotatedabout the optical axis at an very high speed, the observed speckle willbe averaged out. Additionally, such a system reduces the number of scansrequired due to the greater variety of speckle detected.

Since glucose is delivered to the interstitial fluid (IF) in skin viablood, determining the scatter coefficient in the dermis layer of thetissue, where blood vessels are plentiful, provides the closestcorrelation to variations in glucose concentration. Again, an area scanincreases the volume of blood-profused tissue measured.

Area scanning could be achieved by a pair of rotating prisms thatcontinuously move a sample beam spot over a circular area of the tissuesurface. Advantageously, the spot would move a minimum of one spotdiameter for each depth scan. Thus if the beam spot size is 12 micronsand the depth scan is at a rate of 20 Hz, then the spot shouldadvantageously be moved at a minimum rate of 240 microns per second andpreferably much faster.

Spot diameters are typically in the range from about 10 microns to 100microns and preferably 20 microns and higher.

The minimum area of the scan is defined by the number of spot diametersneeded to move at the minimum depth scan rate. For the 12 micron spotand 20 Hz depth scan, the minimum area that would need scanning is about2200 square microns, corresponding to a circular area of about 500micron diameter. More preferably the system would be designed to coveran area corresponding to a diameter of 500 microns to 10,000 microns.

For speckle reduction using the boiling speckle method of noisereduction, the multiple spots would need to be moved quite rapidly. Thespot should move at a minimum of one spot diameter in the integrationtime of the receiver. For individual spot sizes of about 10 microns andan integration time of about 4 microseconds, the spots would need tomove at a minimum of 2.4×10⁵ microns/sec.

The light sources 406 can be light emitting diodes (LED) or superluminescent diodes (SLEDs), both of which are semiconductor based lightemitters whose wavelengths can be chosen to give the best contrastbetween absorption and scatter of blood and other biologicalconstituents, such as water. Typically these wavelengths are in thered/near infrared (RNIR) region of the spectrum, however, longer andshorter wavelengths can be used for enhanced sensitivity. For theglucose measurements, two or more light sources are advantageous and canshare the same optical paths through the interferometer.

One of the wavelengths can be chosen to have minimum absorption comparedto the scattering coefficient for water and blood constituents. If theother wavelength is chosen to have peak absorption for certainbiological constituents, then the difference in light attenuationbetween the two wavelengths can indicate the position in depth of arelevant structure, such as a blood vessel. Light from the twowavelengths is differently absorbed by the different constituents. Thisdifferential absorption differentially reduces the intensity of thescattered (reflected) light. Light reflected off the cellular membraneis partially absorbed by the respective constituent for that wavelength.Where the term “light is reflected from the blood” is used, it isunderstood to refer to light reflected from the cells in and around theblood vessels, and the constituent in the blood absorbs some of thelight according to the specific wavelength and glucose level of theblood. These differences in the scattering and absorption propertiesprovide for an optimal correlation between the scattered signal andblood glucose data.

One exemplary application is a first wavelength of about 1310 nm and asecond wavelength of about 820-960 nm. A first wavelength of 1310 nm ischosen because the scattering properties of water and blood and bloodconstituents is at a maximum compared to the absorption properties ofthese fluids. The second wavelength, 820-960 nm, is chosen because theabsorption of light is very high in the presence of hemoglobin, a bloodconstituent, (compared to the first wavelength). If the signal of thesecond wavelength were to experience a rapid decrease at a particulardepth in the interaction volume, this rapid decrease would indicate thepresence of hemoglobin, and hence, the location of blood-profusedtissue. It would thus indicate an optimal slope region for thescattering data of the first wavelength to be related to the glucoseconcentration.

A second example would be a first wavelength of about 1310 nm and asecond wavelength of about 1450 nm. At this second wavelength, thescattering coefficients for blood and water are similar to those of thefirst wavelength. However, the absorption coefficient for water at thissecond wavelength is exponentially larger than that of the firstwavelength. Thus, a differential measurement between these twowavelengths indicates changes in the hydration level of the tissue. Suchchanges can then be used to indicate an optimal slope region formeasuring blood glucose. However, the use of these two specificwavelengths provides an additional benefit of sensor calibration. As thehydration level in the dermis layer varies, the scattering coefficientof the first wavelength may drift, even though the glucose concentrationremains static. Thus, by measuring the skin hydration using the secondwavelength, this drift can be compensated for and the OCT sensor canmaintain calibration.

An exemplary analysis of the data to determine glucose level essentiallyinvolves the following steps:

-   1. Expressing the reflected intensity I_(r) and the incident    intensity I_(o) as logarithms, e.g. Ln(I_(r)), Ln(I_(o)).-   2. Plotting the logarithmic data in accordance with the scattering    equation. Since the data at the 1310 nanometer wavelength is    dominated by scattering (with minimal absorption) the logarithmic    scattering equation can be approximated by    Ln(I _(r))=L _(n)(I _(o))−(μ_(t))(d)    Where μ_(t) is the scattering coefficient and d is the depth of the    scan.-   3. Determining μ_(t) by regression analysis. This can be achieved by    a linear regression through the logarithmic data to determine the    best fit slope (μ_(t)). Since, however, the glucose concentration is    most accurately read in blood/IF, the regression should analysis is    preferably selectively applied to those data points whose depth (or    two-wavelength scattering characteristics) are indicative of    blood/IF. The glucose concentration, as noted in connection with    FIG. 2 b above, at such is regions strongly correlated with the    scattering coefficient μ_(t) and can readily calibrated to it.    This, or comparable algorithms can be readily implemented in digital    processors by those skilled in the art to provide a rapid,    non-invasive readout of glucose level.

It is understood that the above-described embodiments are illustrativeof only a few of the many possible specific embodiments, which canrepresent applications of the invention. Numerous and varied otherarrangements can be made by those skilled in the art without departingfrom the spirit and scope of the invention.

1-35. (canceled)
 36. A method of monitoring blood glucose concentrationwithin a biological tissue comprising the steps of: providing lighthaving scattering properties sensitive to the blood glucoseconcentration within the tissue; scanning light over a two-dimensionalarea of the tissue; causing a reflected portion of the scanned light tointerfere with a reference beam having an adjustable phase path;obtaining interferometric signals from at least one tissue depth; andprocessing the interferometric signals to determine the blood glucoseconcentration within the tissue.
 37. The method of claim 36 wherein thestep of providing light comprises providing light from one or moresources that is split into a sample beam and the reference beam, andfurther wherein the step of scanning comprises continuously scanning thesample beam over the two-dimensional area, the method furthercomprising: varying the phase of the reference beam before the step ofcausing the reflected portion of the scanned light to interfere.
 38. Themethod of claim 37 further comprising: adjusting the reference beam toconstructively interfere with reflected scanned light from a selecteddepth within the tissue.
 39. The method of claim 36 wherein thetwo-dimensional area is aligned with a skin surface.
 40. The method ofclaim 36 wherein the step of scanning the light over a two-dimensionalarea comprises scanning the biological tissue to reduce speckle.
 41. Themethod of claim 40 further comprising: scanning multiple depths withinthe tissue to reduce speckle.
 42. The method of claim 36 wherein thelight comprises low coherence light.
 43. The method of claim 36 whereinthe light comprises at least two different wavelengths, a firstwavelength having measurably different scattering properties forglucose-containing tissue or indicators of such tissue relative to asecond wavelength.
 44. The method of claim 43 wherein the at least twodifferent wavelengths comprise at least one wavelength at about 1310 nm,at about 1450 nm, and in a range from about 820 nm to about 960 nm. 45.The method of claim 36 wherein the light comprises at least one ofinfrared light and near infrared light.
 46. The method of claim 36wherein the processed measurements comprise non-imaged tissuemeasurements.
 47. The method of claim 36 wherein the reflected portionof the scanned light comprises light reflected from blood-perfusedtissue.
 48. The method of claim 36 further comprises: measuring changesin a hydration level of the tissue.
 49. The method of claim 48 whereinthe step of processing the light measurements comprises compensating forchanges in hydration level of the tissue.
 50. An apparatus for measuringthe blood glucose concentration of a human or animal subject comprising:at least one light source for providing sample beam and reference beam,the at least one light source providing at least one wavelength oflight, the at least one light source configured to direct the samplebeam to tissue of the subject; a phase shifter configured to adjust thephase path of the reference beam; a scanner to scan the sample beam overa two-dimensional area of the subject's tissue; an interferometer toconstructively interfere reflected sample beam from the subject tissuewith the reference beam; at least one photodetector to detect theinterfered light; and a processor to calculate glucose concentrationfrom the detected interfered light.
 51. The apparatus of claim 50wherein the at least one light source comprises a low coherence source.52. The apparatus of claim 50, wherein the interferometer comprises alow coherence interferometer.
 53. The apparatus of claim 50 furthercomprising: a beam splitter to split light from the at least one lightsource into the sample beam and the reference beam, wherein the phaseshifter is configured to direct the reference beam to interfere with thereflected sample beam.
 54. The apparatus of claim 53 wherein the beamsplitter comprises a 2×2 optical splitter.
 55. The apparatus of claim 50wherein the phase shifter comprises an adjustable reflector.
 56. Theapparatus of claim 55 wherein the reflector is configured to scanmultiple depths within the subject tissue.
 57. The apparatus of claim 50wherein the photodetector comprises a photodiode.
 58. The apparatus ofclaim 50 wherein at least one light source comprises a light emittingdiode (LED).
 59. The apparatus of claim 50 wherein at least one lightsource comprises a superluminescent diode (SLED).
 60. The apparatus ofclaim 50 wherein the scanner comprises a movable lenslet array which canbe translated or rotated rapidly to reduce speckle.
 61. The apparatus ofclaim 50 wherein the scanner comprises a movable diffractive opticalelement which can be translated or rotated rapidly to reduce speckle.62. The apparatus of claim 50 wherein the scanner is configured to scanthe sample light such that the two-dimensional area is aligned with askin surface.
 63. The apparatus of claim 50 wherein the scanner isconfigured to scan the sample light to reduce speckle.
 64. The apparatusof claim 50 wherein the apparatus is configured to produce non-imagedinterfered light.